A monolayer is a one-layer-thick film of at least one amphiphilic compound or composition that forms at the air/water interface of an aqueous solution. Molecules in the monolayer are aligned in the same orientation, with the hydrophobic domain facing the air and the hydrophilic domain facing the aqueous solution. Compression of the monolayer results in the formation of an ordered, two-dimensional solid that may be transferred to a substrate by passing the substrate through the monolayer. A monolayer that has been transferred to a substrate is termed a Langmuir-Blodgett film, or LB film. For reviews of Langmuir-Blodgett technology, see Gaines, G. L. Jr. (1966) Insoluble Monolayers at Liquid-Gas Interfaces, Interscience, New York; Zasadzinski et al. (1994) Science 263:1726-1733; Ullman (1991) An Introduction to Ultrathin Organic Films, Academic Press, Boston, Mass.; and Roberts (1990) Langmuir-Blodgett Films, Plenum, N.Y.; the contents of which are incorporated herein by reference.
Monolayers are typically composed of organic molecules such as lipids, fatty acids and fatty acid derivatives, fat-soluble vitamins, cholesterol, chlorophyll, valinomycin and synthetic polymers such as polyvinyl acetate and polymethyl methacrylate, but may also be formed by many other amphiphilic compounds. LB films may be used to detect a molecule that binds to or reacts with a compound of interest that comprises the monolayer or has been incorporated into the monolayer.
Sensing systems employing LB films include electrochemical devices using ion-sensitive field effect transistors, absorption or fluorescence based optical devices, and piezoelectric crystals. For example, LB films of valinomycin have been used to detect a specific interaction of potassium that results in a conformational change that is detectable by infrared spectroscopy (Pathirana et al. (1992) Langmuir 8:1984-1987).
Monolayers incorporating fluorescein lipids have been deposited on quartz crystal microbalances and used to detect specific anti-fluorescyl monoclonal antibodies in solution (Ebato et al. (1994) Analytical Chem. 66:1683-1689). In contrast, detection of antigens by piezoelectric crystals coated with LB films incorporating antibodies has met with limited success. In these systems, non-specific binding of molecules to the LB film prevents accurate measurement of antigen.
Previous methods for forming LB films require dissolution of the compounds to be formed into a monolayer in a volatile organic solvent. The organic solvent forms a separate phase from the aqueous solution and functions to prevent dissolution of the monolayer components in the aqueous phase. After spreading the mixture at the air-liquid interface of the aqueous solution, the solvent is allowed to evaporate, leaving a monolayer at the interface. Unfortunately, the organic solvent often damages the monolayer components and leaves an undesirable residue. LB films formed from such monolayers may have unacceptable levels of nonspecific binding. Such non-specific binding, which is non-saturable, hampers quantitative measurement of specific binding. Our previous invention (U.S. patent application Ser. No. 09/452,968, filed Dec. 2, 1999) overcame such problems by providing a method for forming monolayers that does not require the use of an organic solvent.
Efficient detection using a biosensor device requires: (1) high surface density of functional molecules; (2) high specificity of interactions and the absence of non-specific binding; (3) accessibility of interacting partners; and (4) stability of the sensing system. From a practical standpoint, the most important feature of any biosensor is the dynamic response-time curve of the sensor. When a biosensor is exposed to a specific ligand, the dynamic output signal as a function of time represent the binding process. The total binding (T) includes a non-saturable constituent of non-specific binding (NSB) and a saturable constituent of specific binding (SB). The SB constituent is saturated when the interaction of analytical or diagnostic probe attached to the sensor, for example, a peptide probe, and a target in solution (ligand) reaches a steady-state level. The ability of the probe-ligand system to achieve a steady state level is extremely important for measuring the target ligand concentration in the solution being analyzed.
Unfortunately, extra ligands may be bound to the sensor by the nonspecific interaction with the probe-supporting components. When this occurs, the sensor output corresponding to the steady-state level of specific binding is masked by the increasing contribution of non-specific binding. In practice, to relate concentrations of ligands in a solution being analyzed to sensor output, various variables must be controlled and/or known, such as the volume of liquid, the flow rate of liquid, and the time of exposure. In contrast, when non-specific binding is low, the steady-state output corresponds to a specific ligand concentration. Thus, for optimal performance of sensor devices, surface density and purity of probes must be high and non-specific binding must be minimized.
A critical step in the production of a biosensor is the immobilization of the probe to the surface of the biosensor. Previous methods included a combined Langmuir-Blodgett (LB)/molecular assembly method (Samoylov et al. (2002) Biomolecular Engineering 18: 269-272; Samoylov et al. (2002) J. Mol. Recognit. 15: 197-203). This method involves LB film deposition, which is known in the art and described in references such as Sukhorukov et al. (1996) Biosens. Bioelectron. 11: 913-922; Petty (1991) J. Biomed. Eng. 13: 209-214; Pathirana et al. (1992) J. Am. Chem. Soc. 114: 1404-1405; Pathirana et al. (1992) Langmuir 8: 1984-1987; Pathirana et al. (1996) Supramolecular Sci. 3: 149-154; Pathirana et al. (1998) Langmuir 14: 679-682; Vodyanoy et al. (1994) Langmuir 10: 1354-1357. In some methods, phage-derived probes are directly adsorbed to the sensor device to create a biosensor.
Biosensors previously reported in the literature are somewhat limited because the reported devices have low sensitivity, limited longevity, and/or long response times. Decker et al. ((2000) J. Immunol. Methods 233:159-165) reported that more than 90 minutes were needed to measure phage binding by peptide fragments immobilized by biotin/streptavidin coupling. Hengerer et al. ((1999) Biotechniques 26: 956-60, 962, 964) reported binding of phage antibodies to antigen immobilized on a quartz crystal microbalance with a time constant of about 100 min. These long response times are not compatible with rapid screening and make large-scale screening unwieldy. Therefore, there remains a need for a biosensor which can rapidly detect specific proteins. In addition, reported biosensors generally suffer from disadvantages such as low specificity and low affinity.
Some biosensor platforms utilize antibodies as the binding element. For example, U.S. Pat. No. 5,922,183 teaches the use of thin film composites of metal oxides and antibodies for amperometric and potentiometric sensing. Porous silicon biosensors are described for use with antibodies in U.S. Pat. No. 5,874,047. A patterned multiple antibody substrate for use in biosensors or immunosensors was prepared by adsorbing specific antibodies at the sites in U.S. Pat. No. 5,858,801. U.S. Pat. No. 5,039,611 teaches the use of monoclonal antibodies to superficial papillary bladder tumor cells in an ELISA-type format. See also, copending U.S. application Ser. No. 09/452,968, filed Dec. 2, 1999.
Antibody-based sensors represent an improvement over previously-used sensors in several ways, and can exhibit improved specificity and affinity (see, e.g., Ziegler et al. (1998) Biosensors & Bioelectronics 13: 539-571. However, antibody-based sensors have several disadvantages which restrict their usefulness, including high cost and short longevity or inability to perform in various environmental or field test conditions. Moreover, the quality of antibodies can vary with different production variables, such as the animal used to produce the antibodies. Another disadvantage of antibodies is that it may take months to generate the desired antibodies for use in an antibody-based sensor.
The threat of bioterrorism highlights the need for specific, accurate sensors that are rapidly prepared. At present, the earliest recognition of and response to a bioterrorist attack with Bacillus anthracis (anthrax) spores may be based on clinical manifestations of anthrax and laboratory culture tests, which require days to complete (see, e.g., Inglesby et al. (1999) JAMA 281: 1735-45). Thus, a need exists for specific, accurate biosensors that are rapidly prepared.
Phage-based biosensors have been previously developed. See, e.g., our application Ser. No. 10/289,725, filed Nov. 7, 2002. Typically, a biotinylated monolayer is deposited onto the surface of the sensor device. Following this step, a phage layer may be added using non-LB, molecular self-assembly of a phage layer using biotin/streptavidin coupling. See Furich et al. (1996) SPIE 2928: 220-225 and Volker and Siegmund (1997) EXS 80: 175-191.
Monolayer coverage provides a proximate binding of analyte to sensor surface and therefore works better for sensors in which the short distance between sensor surface and analyte binding site is critical for generation of measurable signal such as acoustic wave or surface plasmon resonance sensors. The present invention provides monolayers of superior purity that provide higher specificity and lower non-specific binding with less manipulation and effort than previous methods, leading to more economical, rapid, and accurate detection of ligands.